This invention relates to optical imaging
Relevant Patents
U.S. Pat. No. 5,076,672 All-optical switch apparatus using a nonlinear etalon Tsuda, et al.
U.S. Pat. No. 5,275,168 Time-gated imaging through dense-scattering materials using stimulated Raman amplification, Reintjes, J et al.
U.S. Pat. No. 5,291,267 Optical low-coherence reflectometry using optical amplification Sorin et al
U.S. Pat. No. 5,299,170 Apparatus for measuring pulse width with two photon absorption medium,. Shibata et al
U.S. Pat. No. 5,321,501 Method and apparatus for optical imaging with means for controlling the longitudinal range of the sample, Swanson E. et al.
U.S. Pat. No. 5,418,797 Time gated imaging through scattering material using polarization and stimulated Raman amplification, Bashkansky et al
U.S. Pat. No. 5,489,984 Differential ranging measurement system and method utilizing ultrashort pulses Hariharan, et al.
U.S. Pat. No. 5,491,524 Optical coherence tomography corneal mapping apparatus Hellmuth, T et al
U.S. Pat. No. 5,549,114 Short coherence length, doppler velocimetry system Petersen, C et al.
U.S. Pat. No. 5,530,544 Method and apparatus for measuring the intensity and phase of one or more ultrashort light pulses and for measuring optical properties of materials Trebino, R et al.
U.S. Pat. No. 5,570,182 Method for detection of dental caries and periodontal disease using optical imaging, Nathel; H et al.
U.S. Pat. No. 5,585,913 Ultrashort pulsewidth laser ranging system employing a time gate producing an autocorrelation and method therefore Hariharan, A et al.
U.S. Pat. No. 5,648,866 Optimized achromatic phase-matching system and method Trebino et al.
U.S. Pat. No. 5,862,287 Apparatus and method for delivery of dispersion compensated ultrashort optical pulses with high peak power Stock et al.
U.S. Pat. No. 5,936,732 Apparatus and method for characterizing ultrafast polarization varying optical pulses Smirl et al.
U.S. Pat. No. 5,920,373 Method and apparatus for determining optical characteristics of a cornea Bille, J
U.S. Pat. No. 5,920,390 Fiberoptic interferometer and associated method for analyzing tissue Farahi, et al.
U.S. Pat. No. 5,975,697 Optical mapping apparatus with adjustable depth resolution Podoleanu, A et al.
U.S. Pat. No. 5,994,690 Image enhancement in optical coherence tomography using deconvolution Kulkarni, M et al.
U.S. Pat. No. 6,002,480 Depth-resolved spectroscopic optical coherence tomography Izatt; J. et al
U.S. Pat. No. 6,006,128 Doppler flow imaging using optical coherence tomography Izatt; J. et al
U.S. Pat. No. 6,008,899 Apparatus and method for optical pulse measurement Trebino; R et al.
U.S. Pat. No. 6,023,057 Device for determining the phase errors of electromagnetic waves Gaffard et al.
U.S. Pat. No. 6,053,613 Optical coherence tomography with new interferometer Wei Jay et al.
U.S. Pat. No. 6,095,651 Method and apparatus for improving vision and the resolution of retinal images Williams, D et al
U.S. Pat. No. 6,111,645 Grating based phase control optical delay line Tearney, et al.
U.S. Pat. No. 6,134,003 Method and apparatus for performing optical measurements using a fiber optic imaging guidewire, catheter or endoscope Teamey et al.
U.S. Pat. No. 6,199,986 Rapid, automatic measurement of the eye""s wave aberration Williams, D et al
U.S. Pat. No. 6,191,862 Methods and apparatus for high speed longitudinal scanning in imaging systems Swanson; A. et al.
U.S. Pat. No. 6,195,617 B I Autocorrelation of ultrashort electromagnetic pulses Reid et al.
U.S. Pat. No. 6,201,608 Method and apparatus for measuring optical reflectivity and imaging through a scattering medium Mandella et al.,
U.S. Pat. No. 6,226,112 Optical Time-division-multiplex system by Denk, et al.
U.S. Pat. No. 6,249,630 Apparatus and method for delivery of dispersion-compensated ultrashort optical pulses with high peak power Stock et al
U.S. Pat. No. 6,256,102 Dual-beam low-coherence interferometer with improved signal-to-noise ratio Dogariu A et al.
U.S. Pat. No. 6,291,824 Apparatus and method for high-bandwidth optical tomography Battarbee, et al
U.S. Pat. No. 6,356,693 semiconductor optical pulse compression waveguide, Shimazu et al.
Other Publications
xe2x80x9cImaging Objects Hidden in a Highly Scattering Media Using Femtosecond Second-Harmonic-Generation Cross-Correlation Time Gatingxe2x80x9d, Yoo et al, Optics Letters, July 1991, pp. 1019-1021. Jenkins and White, fundamental of Optics, McGraw-Hill, 1957
It is well known that a Michelson Interferometer enables to make precise distance and incremental displacement measurements by observing the fringes formed by the interference of coherent light waves. The interference between light waves that have traveled along different pathways is limited by the coherence length of the light source. As long as the different pathways differ by less than the coherence length of the source, interference will result in formation of fringes.
Optical Coherent Tomography (OCT) makes use of a Michelson interferometer to image the topography of the layers behind the surface of a tissue by scanning xe2x80x9csame-depthxe2x80x9d layers. This is achieved by precise balancing of the legs of the interferometer, so that the depth information is obtained by observing the interference fringes when the two legs of the interferometer are within the coherent length of the illuminating light source. Changing the length of one of the paths enables to focus on a layer at a depth that differs by the length changed. However as fringes of equal intensity are obtained with widely differing path lengths, for as long as the interfering light waves are coherent, light sources with short coherence lengths such as superluminescent diodes are used, so as to minimize this ambiguity. This setup greatly facilitates the calibration of the interferometer as no interference fringes are obtained when the path lengths between the two legs of the interferometer differ by more than the coherence length.
However, it is important to realize that the fringes observed with any light source, originate from the interference of light coming from many oscillators which emit light randomly and non-coherently one from the other. Low coherence length sources are limited in resolution by the randomness of the coherence lengths of the different oscillators and the FWHM of the group of fringes is what determines the xe2x80x9cpath-length differencexe2x80x9d resolution and not the FWHM of a single fringe. It is also important to realize that the non-coherence among the various oscillators, also manifests itself in a high uniform background over which the fringe pattern is observed, thus the SNR obtained with low coherence length superluminescent diodes is much worse than the SNR of a fringe pattern obtained with highly coherent sources.
The conventional Optical Coherent Tomography (OCT) technique, (see for example U.S. Pat. No. 5,321,501, Method and apparatus for optical imaging with means for controlling the longitudinal range of the sample, Swanson E. et al.) uses a low coherence light source, to minimize the spread of the fringe pattern and thus increase the xe2x80x9cpath-length differencexe2x80x9d precision.
OCT is constrained by the need to sequentially adjust the depth of the imaged layer by incrementally changing one of the legs of the Michelson interferometer, either mechanically with a retroreflector, by stretching the optical fiber with a piezoelectric motor or by a combination of an acousto-optic deflector, a grating and a mirror (see U.S. Pat. No. 6,111,645 Grating based phase control optical delay line Tearney, et al.). In spite of all the heroic efforts, it takes xcx9c100 microseconds to change the delay, position and balance the interferometer onto a new layer.
OCT is also limited by xe2x80x9cspecklesxe2x80x9d, a background generated by the interference with the coherent multiple back-scattered light, that originates from a spherical volume with a radius equal to the low coherent length of the source.
Ultrafast femtosecond lasers have several important advantages over CW or long-pulse lasers. They permit to achieve high peak power while the average power is relatively low and thus can stimulate nonlinear processes such as second harmonic generation, and amplification. through Stimulated Raman Scattering.
Time gating of Raman amplified signals transmitted through a light diffusing medium in order to locate a strongly absorbing region within such medium, has been demonstrated by Reintjes, et al (see U.S. Pat. No. 5,275,168 Time-gated imaging through dense-scattering materials using stimulated Raman amplification.). Properly adjusting the time delays enable to amplify only the early arriving non-scattered photons, while leaving the multiple scattered diffuse light non-amplified.
U.S. Pat. No. 5,418,797 Time gated imaging through scattering material using polarization and stimulated raman amplification by Bashkansky et al, teaches how to reject the diffuse light by making use of the different polarizations of the diffuse and the non-scattered beams. Note that transmission and reflection geometries are totally different. In a reflection geometry, there are no non-scattered photons, and photons scattered backwards from the different layers, exhibit a continuous distribution in their time-of-flight.
Non-linear crystal such as KDP, KTP or BBO are used in commercially available autocorrelators to establish optical coincidence between two coherent branches of short pulses fed co-linearly into them. The two coherent waves generate a Second Harmonic Generation (SHG) wavelength at half the wavelength, during the spatially overlapping time period and may be detected by a photodetector. The pulse shape is determined by delaying one of the two coherent waves and measuring the intensity at the output of the non-linear crystal. Alternatively measuring the intensities of the spectral content of the pulse as a function of delay will give both its intensity shape and phase.
A narrow temporal width is associated with a wide spectral distribution and thus a single femtosecond laser may be used for multiwavelength excitation of the sample.
U.S. Pat. No. 5,585,913 Ultrashort pulsewidth laser ranging system employing a time gate producing an autocorrelation and method therefore by Hariharan, A et al. teaches a method to measure the topography of a surface by correlating the illuminating femtosecond pulse and the radiation reflected from the examined surface using an SHG (Second Harmonic Generation) crystal.
U.S. Pat. No. 6,249,630 xe2x80x9cApparatus and method for delivery of dispersion-compensated ultrashort optical pulses with high peak powerxe2x80x9d by Stock et al. teaches to stretch the width of optical pulses in order to reduce the peak power transmitted through a fiber and then recompressing it before delivering it to the target.
It is well known that scattering changes the polarization of the scattered wave and therefore using proper polarization analyzers, single scattered photons may be separated from multiple scattered ones.
The speed of light decreases in direct proportion to the increase of the refraction index of the medium in which it propagates. Thus a wide beam passing through a medium whose refraction index changes across the width of the beam will have its different components moving ahead or lagging behind. Thus GRadient INdexed materials that have gradually changing refraction indexes may be used to temporally reshape the wavefront and compensate for time dispersion.
The invention is an imaging device consisting in a high resolution time-of-flight measurement, of a temporally narrow, but spectrally wide, light beam generated by a femtosecond laser source, after being back-scattered by a relatively thick object, whose layers are to be characterized. Those characteristics include, absorbing, elastic and inelastic scattering cross sections, including intensity, polarization, spectral content and the angular distribution of the beam scattered from the various layers penetrated by the illuminating beam. The impinging beam invariably penetrates a certain depth of the object and sometimes traverses or is scattered by it, the degree of which depends on the beam""s wavelength, intensity, angle of incidence and the composition of the scattering medium, that collectively determine the degree of scattering and absorption cross sections.
Contrary to prior art methods that measure one distance at a time, it is a purpose of this invention to collect the data pertaining to the characteristics listed above from all the voxels along the axis of penetration, during a single femtosecond pulse of the illuminating laser, process and store such data during the period between two consecutive pulses of the high repetition rate femtosecond laser.
The time of flight of the back scattered photons and consequently their depth coordinate is determined by measuring their coincidence with the illuminating ultrashort pulse. Such coincidence is established by a time-gate that may be a non-linear medium such as an SHG (Second Harmonic Generation) medium, a Raman-active medium, a non-linear fiber coupler, or a phase-sensitive interferometer. Obviously the speed of the time-gate determines the time-of-flight accuracy and the ability to temporally differentiate between photons back-scattered from consecutive layers, thus determining the degree of characterization of the different layers.
The temporally narrow illuminating beam, when temporally stretched and wavelength filtered will cause its transmitted spectral components to arrive at the scattering body sequentially and then back-scattered. In this case the temporal separation of the spectral components each from the other, has to be larger than the temporal spread of the illuminating pulse caused by back-scattering from the different layers, but smaller than the repetition rate of the femtosecond laser. For example a 10 fs pulsewidth of a f=100 MHz femtosecond laser, which illuminates the target every (1/f)=10 nsec., will be temporally spread to xcex94TL=5 psec after being back scattered from a L=1 mm thick tissue; thus the temporal separation between consecutive wavelengths has to be larger than xcex94TL=5 psec, say xcex94txcex=10 psec. In this case, the total number of wavelengths that can be inserted between two consecutive pulses of the femtolaser is 1/fxcex94txcex=103. When the back-scattering is elastic, the wavelength of the back-scattered photons will not change and in addition to their time of flight sorting, they may be classified in real time according to the wavelength of the illuminating beam by passing them through a passive component such as a grating.
The wavelength of the illuminating beam may also be changed by physically inserting an appropriate interference filter on the path of the temporally stretched femtosecond pulse, using a fast translating motor.
Measuring the spectral back-scattering intensity of a body, while rapidly scanning it, enables to dynamically map regions and structures exhibiting different absorption cross sections. Thus for example the web of vessels transporting the blood may be mapped and the state of oxygenation of the surrounding cells, as a function of the systolic or diastolic pressure may be recorded.
Spectral and temporal cross-correlation between the impinging and scattered beams enables to extract the change of phase, enabling to map same-phase biological tissue structures as indicative of their equivalence.
The extremely narrow pulses having high instantaneous power, result in a high signal/noise ratio and enable to collect all the needed information for a single spot, during a single femtosecond pulse, obviating the need to integrate the signal for a relatively long time, a process usually necessary in order to improve the Signal-to-Noise ration (SNR).
The simultaneous collection of all the time-of-flight data of the photons back-scattered from the different layers, is made possible by a chain of linked AND time-gates equivalent to an xe2x80x9coptical serial-to-parallel converterxe2x80x9d that converts the inherently serial xe2x80x9ctime-of-flightxe2x80x9d information, to parallel optical signals, on the fly, each signal representing the intensity of the back-scattered photons for a different time-of-flight. This method reduces the total volumetric imaging time by a factor equal to the number of layers to be imaged, in fact opening up applications that are not practical to do with the prior-art methods, such as OCT , Confocal microscopy or time-of-flight ranging.
It has to be realized that collecting the back-scattered photons from one layer at a time, as is done by prior-art methods, not only takes more time but is also wasteful from the point of view of photon statistics and signal-to-noise ratio (SNR), given the minimal time of illumination required in all dynamic applications where the object is moving. The impinging beam always penetrates the maximal allowable depth determined by the physics of the interaction, and is scattered by all the interim layers. Limiting the collection of back-scattered photons to the surface or one layer, and rejecting the photons back-scattered from all the other layers, is a tremendous waste, a waste that increases with improvement of the axial resolution.
To illustrate our argument numerically, if 100 layers are imaged sequentially, one at a time, 99% of the information is lost and given a fixed total time of imaging, the SNR will be (100)1/2=10 times worse. In Ophthalmology for example, where damage to the retina has to be avoided and therefore the illuminating intensity limited, throwing away 99% of the information leads to unsatisfactory diagnostic images.
In addition to their precisely determined time-of-arrival, scattered photons may also be sorted according to their state of polarization, thus separating, the once back-scattered photons, from double and multiple scattered photons. The extremely narrow illumination in time of one single voxel combined with a narrow time-gate, reduces drastically the multiple scattering. For example if only 1% of the beam is scattered from within the time defined voxel, twice-scattered photons within the same time-voxel constitute (1%)xc3x97(1%)=10xe2x88x924 of the impinging beam or 1% of the single scattered photons. The solid angle to the collecting detector further reduces the portion of double or multiple scattered (more than twice) scattered photons.
The extremely short information collection time per pixel, combined with a high repetition rate source and high speed beam deflectors further enhanced by the ability to collect the information from all the layers simultaneously, result in data collection and characterization of large volumes, in exceptionally short times.
Thus for example data characterizing 1 million voxels (100xc3x97100xc3x97100 pixels), can be collected within 100 microseconds. Such data collection speed, with spatial resolutions of the order of cellular dimensions, enables to follow kinetics of well defined biological structures. The capabilities described above when applied to vascular and arterial high resolution imaging of blood vessels, by applying dual wavelength illumination, enables to follow temporally, the oxygenation kinetics at the cellular level. Such processes may discriminate cancerous growths from normal tissue based on observation of angiogenesis coupled with the existence of hypoxic regions and polarization characteristics as a function of blood flow. The capability to follow blood kinetics at the millisecond time scale combined with cellular spatial resolutions, enables to follow neurological functions. Dynamic imaging of the vasculature and microvessels enable to discern developing aneurysms and follow embolisms, immediately below the surface.
The time-of-flight method may be used to determine the eyeball""s optical aberrations by measuring directly the shape of the light wave emanating from a point on the retina, when this point is illuminated with a narrow light beam. The arrival time of the reflected/back-scattered rays are measured sequentially for a large matrix, within a short time and the phase of each of the rays is calculated by measuring the cross-correlation with the illuminating beam.